Bone grafts are used to provide anatomic mechanical replacement of bone lost due to trauma or removal of cancerous tissue. The source of human bone graft material is primarily limited to the patient's own skeleton (autogenous) or a limited supply of donated bones (allogenous) maintained in bone banks. There is a need for a suitable synthetic source of compatible bone graft material.
Bone is the second most implanted material in humans after blood. There are over 450,000 bone grafts in the United States and over 2.2 million worldwide. Autogenous bone used as an autograft is superior to allogeneic bone, but has the disadvantages of donor site morbidity, prolonged surgery time and limitations in quantity of graft available (especially when large, load bearing cortical segments are required). Collection of the needed material increases the surgical anesthetic time, risk of infection, and intra-operative blood loss. Sufficient quantities and dimensions may not be available or may require multiple donor sites; and damage at the donor site may result in unacceptable patient morbidity. In addition, the quality of the material may be poor if it is necessary to harvest additional bone from a previously used autograft site. In part due to these restrictions, bone allografts are used extensively in human surgery, accounting for the materials used in vast majority of bone grafts.
The primary source of bone allograft material is bone banks. Bone banking was pioneered by the military in the Korean War period due to the need for bone grafts in reconstructive surgery after trauma. The American Association of Tissue Banks (www.aatb.org) was established in 1976 as a not-for-profit, peer group organization to facilitate the provision of tissue on a national basis. AATB publishes standards and technical manuals for collection, processing, storage, and distribution of tissue, and accredits tissue banks as complying with these standards. The majority of available bone allograft material in the United States is obtained and maintained by the limited number (about 60) of AATB approved bone banks.
However, the availability of bone bank material is limited by the possibility of the transmission of infectious diseases, notably HIV/AIDS and hepatitis B and C. The U.S. Food and Drug Administration regulates both the screening of domestic bone bank material and the importation of foreign tissue, for which there may not be proper information on source and donor health. The number of domestic organ donors has been relatively constant at about 5,000-6,000 cadaveric donors per year. At present, the number of patients waiting for donated organs and tissue far exceeds the number of donors each year.
Another option is the xenograft, which is a living or non-living transplant from another species. Non-living, chemically treated xenografts are regularly used as heart valve replacements. Bovine bone grafts are seldom used due to concerns over transmission of viruses and prions (e.g., bovine spongiform encephalitis, “mad cow disease”). Hydroxyapatite derived from marine coral is also considered a xenograft. However, harvesting marine coral is limited by an international treaty. An implant of synthetic bone would avoid the problems inherent in traditional autografts, allografts, and xenografts, but only if the synthetic bone graft material was physically, chemically and functionally compatible with natural bone.
Human bone mass, and thus bone density, and strength decreases with age (osteoporosis), making bones particularly susceptible to fracture in older people. Bone mass decreases because the bone-growing cells (osteoblasts) become less productive at making new bone and repairing microfractures. Fractured hips and collapsed vertebrae are a common result of low bone density in older people. A comparison between normal bone and osteoporotic bone is shown in FIG. 1. FIGS. 1A and 1B show a comparison of photomicrographs between normal bone and osteoporotic bone. FIG. 1A shows a photomicrograph of normal cancellous bone of a 30 year old woman and FIG. 1B shows a photomicrograph of osteoporotic cancellous bone from a 60 year old woman, whether the mineral appears as dark areas, and pores appear as light areas.
Natural bone consists mainly of minerals and collagen fibers, with approximately 69% by weight being hydroxyapatite (HAp, Ca10(PO4)6(OH)2). In practice, the chemical composition of bone mineral is a bit more complex than this formula suggests. The calcium sites are doped to about 1.5% with several mono- or divalent cations (Na, K, Mg, Zn, Fe, Sr, Pb, Ba, Cu, etc.). The hydroxyl and phosphate groups are doped with carbonate ions to about 5% by weight, and the mineral has lower oxygen content. Bone mineral has the generic formula of Ca8.3(PO4)4.3(HPO4,CO3)1.7(OH,CO3)0.3.
Mature bone is either cortical bone or cancellous bone. Cortical bone, which is also known as compact bone, is always on the exterior surrounding the cancellous bone. Cancellous bone, also known as trabecular or spongy bone, develops near the ends of long bones, at the interior of small bones, and between the surfaces of flat bones. Cortical bone consists of several irregular cylindrical units, called Haversian systems or secondary osteons, each consisting of a central Haversian canal containing a neurovascular bundle surrounded by concentric lamellae of bony tissue. The Haversian canals average about 50 μm in diameter, with those closer to the marrow cavity being slightly larger. Within each cavity are one or two capillaries and usually some nerve fibers. Cancellous bone consists of an array of plates and rods of bone tissue (“trabeculae”) and forms an open-celled foam. The trabeculae represent “unrolled” osteons. The difference between compact and spongy bone is the amount of solid matter and the size, shape, and number of spaces in each. In compact bone, the spaces are small and the solid matter is extensive, while the opposite is true in spongy bone.
Synthetic bone is typically made from hydroxyapatite or a combination of hydroxyapatite and tricalcium phosphate (Ca3(PO4)2). These materials can be made with a porous structure, which allows natural bone to grow into the pores and form a good bond. Current methods for the production of hydroxyapatite typically involve sintering at high temperatures. However HAp decomposes at the sintering temperatures, approximately 1000° degrees Celsius. Sintering at lower temperatures produces weak bonds, while sintering at higher temperatures causes low porosity and poor phase purity (due to decomposition). This makes sintering an inadequate method for producing good quality HAp implant material for structural bone replacement.
An ideal synthetic bone graft material would be resorbable so it will be gradually broken down and replaced with new natural bone continuously through normal bone remodeling. The material should be porous and osteoconductive to be populated by bone cells. The physical properties of the synthetic bone graft material should be compatible with natural bone to provide mechanical support while participating in the proper transmission of local stress required by the endogenous control mechanism of bone remodeling and bone growth.
A disadvantage of using metals and nonresorbable ceramics as bone implants based on the phenomenon of stress shielding, which was discovered by Julius Wolff in 1892. Materials implanted into bone will the load that the bone would normally experience. The stress that the natural bone is subjected to is based on the ratio of the elastic modulus of the implant material to that of the natural bone. If the elastic modulus of the implant material is much higher than the elastic modulus of the natural bone, the bone will experience proportionally lower stresses. This will cause the bone to adapt to this lower stress by atrophy. The elastic modulus of several materials, including cortical bone, is listed in Table 1. Most nonresorbable materials have a higher elastic modulus than bone, and thus are a poor choice for synthetic bone graft materials.
TABLE 1Mechanical properties of Bone and Selected materialsDensityElasticMaterial(g/ml)Modulus (GPa)Modulus/DensityGraphite Fiber1.8276153.3Alumina3.934588.5Hydroxyapatite3.227987.2Hardwoods1.310076.9Ivory1.99047.4Quartz2.6510338.9Aluminum2.77025.9Stainless Steel8.0219324.1Titanium5.011422.8Zirconium6.58312.8PMMA (solid)1.1832.5UHMW PE0.9411.1Compact Bone2.1209.5Trabecular1.00.10.1Bone
The combination of hard inorganic and flexible organic components gives natural bone tissue nearly equal resistance to compression and tension. Bone is comparable to cast iron in tensile strength, with only a third of the weight. The breaking stress of bone and cast iron are 235 and 273 MPa, respectively.
A porous biomaterial allows natural bone to grow into the implant. Porosity allows entry by osteoclasts and osteoblasts which leads to osteointegration and vascularization. It has been shown that 37% of pore volumes must be enclosed by interconnections greater than 100 μm for mineralized bone in growth. Larger pores and more interconnectivity would produce more rapid and complete ingrowth; however, a tradeoff between ingrowth and implant strength must be achieved as larger pores greatly decrease the strength of the implant.
Ceramics used for the repair, reconstruction, and replacement of diseased or damaged parts of the body are termed bioceramics. Bioceramics can be designed to be very similar in chemistry and structure to natural bone. Bioceramics can be classified into four categories: nearly inert bioceramics, such as alumina and carbon; surface-active bioceramics, such as Bioglass; resorbable bioceramics, such as calcium sulfate and tricalcium phosphate and composites, such as polymer-ceramic composites and HAp/TCP mixtures. See, generally, Vincenzini, P., ed., High Tech Ceramics (Materials Science Monographs, 38A), Elsevier, Amsterdam, 1987. Much work on synthetic bone graft materials has been on Bioglass and other glass-ceramic composites, polymer-ceramic composites and HAp/TCP calcium phosphate mixtures.
Bioglass, developed by Hench and co-workers in the early 1970s, is a glass designed to bond directly to bone by providing surface reactive silica, calcium, and phosphate groups in an alkaline pH environment. Bioglass 45S5 contains 45 wt % SiO2, 24.5 wt % CaO, 24.5 wt % Na2O, and 6 wt % P2O5. However, the orthopedic applications of Bioglass are limited due to the slow kinetics of surface reaction rates and the corresponding slow development of interfacial bond strength. Bioglass also has mechanical properties common to other glass-ceramic composites: low tensile strength, ductility, and modulus mismatch with natural bone.
The most widely used bioceramics are hydroxyapatite (HAp, Ca10(PO4)6(OH)2) and tricalcium phosphate (TCP, Ca3(PO4)2). Calcium phosphate materials have several advantageous characteristics including a lack of local or system toxicity, a lack of inflammatory or foreign body response, an absence of intervening fibrous tissue between implant and bone, and the ability to become directly bonded to bone.
TCP has been shown to resorb rather quickly, but HAp is the more similar to biological apatite crystals in chemistry and physical properties. Hydroxyapatite dissolves much more slowly than TCP in a variety of fluids. It has been shown that when dense HAp and TCP of similar purity and microstructure were compared, the TCP dissolved 12.3 times faster in acid solutions and 22.3 times faster in basic solutions.
A combination of fast resorbing TCP and much stronger Hap, as can be termed biphasic calcium phosphate (BCP), provides a combination of desirable properties. See Legeros, R., et al., “Biphasic calcium phosphate bioceramics: preparation, properties and applications,” Journal of Materials Science: Materials in Medicine, vol. 14, no. 3, 2003, pp. 201-209. BCP mixtures have been studied as bone substitute materials for dental and orthopedic applications (Daculsi, G., et al., Macroporous calcium phosphate ceramic for long bone surgery in humans and dogs. Clinical and histological study, J. Biomedical Materials Research, 24 1990, 379-396).
Biphasic calcium phosphates are a mixture of HAp and beta-TCP. The bioactivity of biphasic calcium phosphate ceramics can be controlled by altering the HAp/beta-TCP ratio. The dissolution rate of a biphasic calcium phosphate ceramic is dependent on the HAp/beta-TCP ratio: the higher the ratio, the lower the extent of dissolution. Biphasic calcium phosphates are osteoconductive but not osteoinductive.
The tissue response to porous HAp differs from that of dense HAp in that porous HAp allows bone ingrowth, i.e., is osteoconductive. Porosity and interconnectivity determine the amount and type of ingrowth. For implants with a high degree of porosity and interconnectivity, tissue ingrowth begins after 3 or 4 days. After 28 days the ingrowth is complete. The bone-HAp bonding found within the pores is similar to that in dense HAp.
One of the most attractive features of calcium phosphates is their ability to become strongly bonded to living bone. The bond is so strong, that implants cannot be detached from the adjoining bone without fracturing either the implant or the bone (Jarcho, M., Biomaterial Aspects of Calcium Phosphates, Dental Clinics of North America, 30, 1986, 25-47.).
The most common method used to prepare calcium phosphate bioceramics, such as HAp and TCP, involves the use of powders prepared from aqueous solutions of the starting chemicals. These powders are compacted under high pressure (10 to 20,000 psi) and then sintered at between 1000° C. and 1300° C. See Jarcho, 1986. Biphasic calcium phosphate (BCP) is obtained when calcium-deficient biologic or synthetic apatites are sintered at or above 700° C. An apatite is considered calcium deficient when the Ca/P ratio is less than the stoichiometric value of 1.67 for pure calcium hydroxyapatite. However, sintering at high temperatures causes low porosity, closed pores, and poor phase purity (due to decomposition), while sintering at lower temperatures to increase porosity produces weak bonds (Brown, P., et al., Factors Influencing the Formation of Monolithic Hydroxyapatite at Physiological Temperature, in Rusin, R. P., et al. eds., Bioceramics. Materials and Applications II (Ceramic Transactions v63), American Ceramic Society, Westerville, Ohio, 1995, pp. 37-48.) Thus sintering is inadequate for producing HAp implant material for structural bone replacement.
Stoichiometric calcium hydroxyapatite ceramics, especially if it is heated at high temperatures, do not take part in the bone remodeling process due to the loss of carbonate ions at high temperatures. The ideal Ca/P ratio of HAp is 10:6 (1.66667), and the calculated density is 3.219 g/cm3. Calcium phosphate can be crystallized into salts, hydroxyapatite, and beta-whitlockite depending on the Ca/P ratio, presence of water, impurities, and temperature. However, in many cases, more than one structure will exist in the same product. At lower temperatures (<900° C.) and wet environments it is more likely that hydroxyapatite will form. At higher temperatures and in dry atmospheres beta-whitlockite will be formed.
Precipitates of hydroxyapatites can be made from an aqueous solution of Ca(NO3)2 and NaH2PO4. One method uses precipitates that are filtered and dried to form a fine particle powder. After calcination for 3 hours at 900° C., the powder is pressed into a final form and sintered at about 1050° C. to 1200° C. for 3 hours. Macroporosity can be generated by incorporating volatile materials (naphthalene or hydrogen peroxide), heating at a temperature below 200° C., and then sintering.
A good bioceramic should have an interconnected porosity between 55 and 70 percent, and the pore size should range from 150 to 700 μm as in natural bone. An ideal cancellous bone graft substitute would mimic osteon-evacuated cancellous bone and have a thin lattice interconnected by pores of 500-600 micrometers. Since the Haversian systems are about 190-230 micrometers in diameter, an ideal bone graft material for cortical bone regeneration would have an interconnected porous system of similar dimensions.
Polymer-ceramic composites and polymer meshworks have been developed to provide suitable porosity, generally at the expense of desirable mechanical properties. See Yang, et al., The design of scaffolds for use in tissue engineering. Part I. Traditional factors, Tissue Eng. 2001 7(6): 679-689; Pompe, W., et al., Functionally graded materials for biomedical applications, Materials Science and Engineering 2003 A362: 40-60.